A conventional computer tomography (CT) scanner includes an x-ray tube that emits ionizing x-ray radiation that traverses an examination region and a patient therein and illuminates a detector array disposed across the examination region from the x-ray tube. The detector array has included a plurality of single-energy integrating detectors, each including a gadolinium oxysulfide (GOS) or other scintillator array optically coupled to a photo-sensor array. A dual-energy detector has included a first scintillator array having a material (e.g., zinc selenide (ZnSe)) for absorbing lower energy “soft” x-ray photons and another scintillator array having a material (e.g., GOS) for absorbing higher energy “hard” x-ray photons.
The x-ray radiation illuminates the scintillator array, which absorbs the x-ray photons and, in response, emits optical photons indicative of the absorbed x-ray photons. The photo-sensor array detects the optical photons and generates an electrical (current or voltage) signal indicative of the detected optical photons. Inevitably, some residual x-ray photons pass through the scintillator without being absorbed by the scintillator array. A reconstructor reconstructs the photo-sensor output signals and generates volumetric image data indicative of the scanned patient. With dual-energy detectors, images can be generated for each energy level and/or the data can be combined to generate an image similar to a single-energy system.
X-ray photons traversing the patient are attenuated and absorbed by the patient as a function of the radiodensity of the tissue being traversed, and the energy deposited in the patient generally is referred to as deposited or patient dose. Unfortunately, such ionizing radiation can damage cells. A trend in CT has been to reduce patient dose, including for screening asymptomatic patients, and/or imaging younger populations, patients undergoing recurrent scans, etc. However, reducing patient dose adversely affects image noise, which is mainly dominated by Poissonic (“quantum”) noise of the x-ray photons arriving at the detector. Furthermore, the attempt to image with relatively lower dose in conventional CT scanners creates significant excess image noise and artifacts when, for example, the electronic signals generated by the detector are close to the level of the electronic noise.
Almost all scintillators have some optical absorption in addition to the optical emission (due to x-ray absorption). This optical absorption is proportional to the distance that the light photons travel as they are scattered to the photo-detector. As such, increasing the height of a vertical photo-detector (where the photo-detector is mounted to the side of the scintillator relative to the direction of the incident radiation) without increasing x-ray photon absorption can reduce optical photon absorption and, thus, increase the level of the electronic signals generated by the detector.
Another trend in CT has been to increase spatial resolution, while maintaining a predefined coverage area. The resolution can be increased by reducing, given a fixed distance from the x-ray focal spot, the dimensions of the detectors. However, in order to maintain the predetermined coverage area with detectors having a reduced dimension, more detectors are required, which may increase the overall cost of the detection system. Furthermore, in order to maintain a given signal to noise level for an image, the electrical circuitry on a detector should be located as closely as possible to the photo-detector to reduce additional electronic noise. A relatively higher level of performance can be achieved when the electrical circuitry is next to the photo-sensor. However, placing the circuitry in such a location requires either blocking residual the x-ray radiation from reaching the electrical circuitry (e.g., via a radiation shield) or using radiation hardened electrical circuitry, which, respectively, can consume more area and/or increase detector cost.